In the field of medicine, radiation may be used for diagnostic, therapeutic and palliative purposes. For example, the therapeutic use of radiation such as x-rays and y-rays may involve eradicating malignant cells. Conventional radiation treatment systems used for medical treatment, such as linear accelerators that produce high-energy x-rays, utilize a remote radiation source external to the targeted tissue. A beam of radiation is directed at the target area, for example a tumor inside the body of a patient. The x-rays penetrate the patient's body tissue and deliver radiation to the cancer cells, usually seated deep inside the body. This type of treatment is referred to as teletherapy because the radiation source is located at some distance from the target. This treatment suffers from the disadvantage that tissue disposed between the radiation source and the target is exposed to radiation. To reach the cancer cells, the x-rays from an external radiation source must usually penetrate through normal surrounding tissues. Non-cancerous tissues and organs are therefore also damaged by the penetrating x-ray radiation.
An alternative treatment system utilizing a point source of radiation is disclosed in U.S. Pat. No. 5,153,900 issued to Nomikos et al., U.S. Pat. No. 5,369,679 to Sliski et al., U.S. Pat. No. 5,422,926 to Smith et. al., and U.S. Pat. No. 5,428,658 to Oettinger et al., all owned by the assignee of the present application, all of which are hereby incorporated by reference. This system includes a miniaturized, insertable probe capable of producing low power x-ray radiation while positioned within or in proximity to a predetermined region to be irradiated. The probe may be fully or partially implanted into, or surface-mounted onto a desired area within a treatment region of a patient. X-rays are emitted from a nominal, or effective “point” source located within or adjacent to the desired region to be irradiated, so that a desired region is irradiated, while irradiation of other regions are minimized. This type of treatment is referred to as brachytherapy, a word derived from the ancient Greek word for close (“brachy”), because the source is located close to or in some cases within the area receiving treatment.
Brachytherapy offers a significant advantage over teletherapy, because the radiation is applied primarily to treat a predefined tissue volume, without significantly affecting the tissue adjacent to the treated volume. The term brachytherapy is commonly used to describe the use of radioactive isotopes which can be placed directly within or adjacent the target tissue to be treated. Handling and disposing of such radioisotopes, however, may impose considerable hazards to both the handling personnel and the environment. X-ray brachytherapy offers the advantages of brachytherapy, while avoiding the use of radioisotopes.
X-ray brachytherapy treatment generally involves positioning the insertable probe into or adjacent to the tumor or the site where the tumor or a portion of the tumor was removed to treat the tissue adjacent the site with a local boost of radiation. Radiation probes of the type generally disclosed in U.S. Pat. No. 5,153,900 typically include a housing, and a hollow, tubular probe or catheter extending from the housing along an axis and having a target assembly at its distal end. The probe typically encloses an electron source having a thermionic cathode or a photocathode. The electron source also typically includes an accelerating means for establishing an acceleration potential difference between the electron source and the target. The target emits radiation in response to incident electrons from the electron source.
In conventionally heated thermionic cathodes, a filament is resistively heated with a current. This in turn heats the cathode so that electrons are generated by thermionic emission. In a typical conventional x-ray machine, for example, the cathode assembly may consist of a thoriated tungsten coil approximately 2 mm in diameter and 1 to 2 cm in length which, when resistively heated with a current of 4 A or higher, thermionically emits electrons. Thermionic cathodes must be stable against temperature rise under operation, since they may be subject to several thousand degrees centigrade. In a photocathode, a photoemissive substance is irradiated by a LED or a laser source. Typically, a flexible fiber optical cable couples light from the LED or laser source to the photocathode. The laser beam shining down the fiber optic cable activates the photocathode which generates free electrons by the photoelectric effect. Photocathodes may be subjected to several hundred degrees centigrade.
In order to prevent probe failure, it is important that the electron source be heated as efficiently as possible, namely that the electron source reach as high a temperature as possible using as little power as possible. In conventional x-ray tubes, for example, thermal vaporization of the cathode filament is frequently responsible for tube failure. Also, the anode heated to a high temperature can cause degradation of the radiation output. During relatively long exposures from an x-ray source, e.g. during exposures lasting from about 1 to about 3 seconds, the anode temperature may rise sufficiently to cause it to glow brightly, accompanied by localized surface melting and pitting which degrades the radiation output.
While a photocathode avoids such problems, there are difficulties inherent in fabricating the photocathode, because photocathode fabrication should preferably be done in a vacuum. A photocathode must have a sufficient quantum efficiency, where quantum efficiency relates to the number of electrons generated per incident light quantum. The degree of efficiency must be balanced to the intensity of available incident light. For practical substances, with reasonable quantum efficiencies above 10−3, the fabrication of the photocathode should be performed in a vacuum. U.S. Pat. No. 5,428,658, owned by the assignee of the present application and hereby incorporated by reference, discloses an example of such vacuum fabrication.
It is possible to further increase the efficiency of, and reduce the power requirements of, miniaturized therapeutic radiation sources as discussed above, by using a laser rather than an ohmic current, to heat the thermionic cathode. U.S. patent application Ser. No.   (identified by Attorney Docket Nos. PHLL-155 and hereby incorporated by reference)(hereinafter the “PHLL-155” application) discloses a miniature therapeutic radiation source that uses a reduced-power, increased efficiency electron source. The electron source disclosed in the PHLL-155 application has a laser-heated thermionic cathode,.which generates electrons with minimal heat loss, and which does not require a vacuum-fabricated photocathode. The electron source includes a thermionic cathode having an electron emissive surface. The PHLL-155 application discloses using laser energy to heat the electron emissive surface of the thermionic cathode, instead of resistively heating the electron emissive surface of the thermionic cathode. In this way, electrons can be produced in a quantity sufficient to form an electron current necessary for generating therapeutic radiation at the target, while significantly reducing the requisite power requirements for the radiation source.
It is desirable that the surfaces of the thermionic cathodes be heated to as high a temperature as possible, and as rapidly as possible, i.e. that the surfaces be heated as efficiently as possible. Therefore, one way of reducing the power requirements for a therapeutic radiation source, such as the source disclosed in the PHLL-155 application, is to minimize heat loss by the thermionic cathode. Heat loss by laser-heated thermionic cathodes may include 1) heat lost by thermal conduction; and 2) heat loss caused by the portion of incident laser radiation that remains unabsorbed; and 3) heat loss by thermal radiation. One of the features disclosed in the PHLL-155 application are reflector elements. These reflector elements can reflect back to the thermionic cathode incident laser radiation that remained unabsorbed by the electron emissive surface of the thermionic cathode, thereby minimizing heat loss due to unabsorbed incident laser radiation. These reflector elements cannot reduce, however, heat loss that is caused by thermal conduction in the thermionic cathode.
It is an object of this invention to reduce heat loss that is caused by thermal conduction in a laser heated thermionic cathode, thereby further increasing the efficiency of a laser-driven therapeutic radiation source and reducing the power requirements therefor. It is another object of this invention to provide a thermionic cathode for use in a therapeutic radiation source, where the thermionic cathode is shaped and configured so as to reduce heat loss caused by thermal conduction within the cathode.